Cardiac pump with speed adapted for ventricle unloading

ABSTRACT

A blood pump system is implantable in a patient for ventricular support. A pumping chamber has an inlet for receiving blood from a ventricle of the patient. An impeller is received in the pumping chamber. A motor is coupled to the impeller for driving rotation of the impeller. A motor controller is provided for tracking systolic and diastolic phases of a cardiac cycle of the patient and supplying a variable voltage signal to the motor in a variable speed mode to produce a variable impeller speed linked to the cardiac cycle. The impeller speed comprises a ramping up to an elevated speed during the diastolic phase in order to reduce a load on the ventricle at the beginning of the systolic phase.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.15/640,001, filed on Jun. 30, 2017, which is a continuation of U.S.application Ser. No. 13/873,551, filed on Apr. 30, 2013, now U.S. Pat.No. 9,713,663, the disclosures of which are incorporated herein byreference in its entirety, for all purposes, as if fully set forthherein.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

Not Applicable.

BACKGROUND OF THE INVENTION

The present invention relates in general to ventricular support pumpsand controls, and, more specifically, to a ventricular assist device forreducing load applied to a weakened ventricle during the systolic phase.

Many types of circulatory assist devices are available for either shortterm or long term support for patients having cardiovascular disease.For example, a heart pump system known as a left ventricular assistdevice (LVAD) can provide long term patient support with an implantablepump associated with an externally-worn pump control unit and batteries.The LVAD improves circulation throughout the body by assisting the leftside of the heart in pumping blood. One such system is the DuraHeart®LVAS system made by Terumo Heart. Inc., of Ann Arbor, Mich. TheDuraHeart® system employs a centrifugal pump with a magneticallylevitated impeller to pump blood from the left ventricle to the aorta.The impeller can act as a rotor of an electric motor in which a rotatingmagnetic field from a multiphase stator couples with the impeller and isrotated at a speed appropriate to obtain the desired blood flow throughthe pump.

A typical cardiac assist system includes a pumping unit, driveelectronics, microprocessor control unit, and an energy source such asrechargeable batteries and/or an AC power conditioning circuit. Thesystem is implanted during a surgical procedure in which a centrifugalpump is placed in the patient's chest. An inflow conduit is pierced intothe left ventricle to supply blood to the pump. One end of an outflowconduit is mechanically fitted to the pump outlet and the other end issurgically attached to the patient's aorta by anastomosis. Apercutaneous cable connects to the pump, exits the patient through anincision, and connects to the external control unit. An LVAD system maybe used with or without a pacemaker.

A control system for varying pump speed to achieve a target blood flowbased on physiologic conditions is shown in U.S. Pat. No. 7,160,243,issued Jan. 9, 2007, which is incorporated herein by reference in itsentirety. A target blood flow rate may be established based on thepatient's heart rate so that the physiologic demand is met. In one typeof conventional control unit, a constant speed setpoint has beendetermined for the pump motor to achieve the target flow based ondemand. In this type of system, the pump speed is substantially constantwithin an individual cardiac cycle.

Pulsatile pumping systems are also known wherein the pump speed isvaried within the cardiac cycle to more closely mimic natural heartaction. In one example, U.S. Pat. No. 8,096,935 to Sutton et aloscillates the speed of the pump to produce a pulsed pressure. The speedis oscillated synchronously with the natural cardiac cycle such that apump speed is increased during systole (the time of highest flow) anddecreased during diastole (the time of lowest flow).

Whether operated at a constant speed or in a pulsatile manner, it isknown that when desiring to obtain a maximum unloading of a weakenedventricle the average pump speed should be increased as much as possible(so that the pump flow is increased to the point where it capturesalmost the entire cardiac output). Due to flow inertia, however, thepump flow lags the ventricular pressure increase occurring at thebeginning of systole. Therefore, the ventricle contraction still remainsisometric at the beginning of systole (i.e., the pressure inside theventricle resists its contraction). Furthermore, an increased averagepump speed increases the risk of ventricular suction, particularly atthe end of systole when the ventricle could be nearly empty.

SUMMARY OF THE INVENTION

In order to make ventricular contraction easier, the pump speed isincreased before the systolic phase of cardiac cycle. As a result, theintra-ventricular pressure is reduced prior to ventricular contractionallowing a weak ventricle to contract with reduced resistance. In orderto prevent ventricular suction, the pump speed is reduced to before theend of systole when the ventricle is nearly empty.

In one aspect of the invention, a blood pump system is provided forimplanting in a patient for ventricular support. A pumping chamber hasan inlet for receiving blood from a ventricle of the patient. Animpeller is received in the pumping chamber. A motor is coupled to theimpeller for driving rotation of the impeller. A motor controller isprovided for tracking systolic and diastolic phases of a cardiac cycleof the patient and supplying a variable voltage signal to the motor in avariable speed mode to produce a variable impeller speed linked to thecardiac cycle. The impeller speed comprises a ramping up to an elevatedspeed during the diastolic phase in order to reduce a load on theventricle at the beginning of the systolic phase. In some embodiments,the impeller speed also comprises a ramping down to a reduced speedduring the systolic phase to avoid collapse of the ventricle.

The variable speed mode may be comprised of a constant current mode ormay be comprised of a speed control for matching impeller speed to atarget speed in which the target speed ramps up to the elevated speedduring the diastolic phase and ramps down to a reduced speed during thesystolic phase to avoid collapse of the ventricle.

The motor controller may be configurable to provide the variable voltagesignal to the motor in either the above variable speed mode or aconstant speed mode. The constant speed mode maintains a substantiallyconstant speed of the impeller over each respective cardiac cycle. Aselection between the variable speed mode and the constant speed mode isdetermined according to a physiologic capability of the patient. Thisallows for selective therapy during LVAD support. For example,immediately following the implantation when the left ventricle is weak,the pump is set to operate in the constant current mode therebyproviding a greater level of ventricle unloading. With the patient'srecovery, the pump may be set to operate in the constant speed mode,promoting higher flow pulsatility and a more natural physiologicresponse to the patient's activities.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a diagram of a circulatory assist system as one example of animplantable pump employing the present invention.

FIG. 2 is an exploded, perspective view of a portion of a centrifugalpump of a type that can be used in the present invention.

FIG. 3 is a diagram showing a prior art variable speed profilesynchronized with a cardiac cycle.

FIG. 4 is a diagram showing a variable speed profile of the presentinvention.

FIG. 5 is a diagram showing a variable speed profile in greater detail.

FIG. 6 is a block diagram showing a pump motor and control system of thepresent invention.

FIG. 7 is a diagram showing pump load and speed according to anembodiment using a constant current mode.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Referring to FIG. 1, a patient 10 is shown in fragmentary frontelevational view. Surgically implanted either into the patient'sabdominal cavity or pericardium 11 is the pumping unit 12 of aventricular assist device. An inflow conduit (on the hidden side of unit12) pierces the heart to convey blood from the patient's left ventricleinto pumping unit 12. An outflow conduit 13 conveys blood from pumpingunit 12 to the patient's aorta. A percutaneous power cable 14 extendsfrom pumping unit 12 outwardly of the patient's body via an incision toa compact control unit 15 worn by patient 10. Control unit 15 is poweredby a main battery pack 16 and/or an external AC power supply and aninternal backup battery. Control unit 15 includes a commutator circuitfor driving a motor within pumping unit 12.

FIG. 2 shows a centrifugal pump unit 20 having an impeller 21 and a pumphousing having upper and lower halves 22 a and 22 b. Impeller 21 isdisposed within a pumping chamber 23 over a hub 24. Impeller 21 includesa first plate or disc 25 and a second plate or disc 27 sandwiched over aplurality of vanes 26. Second disc 27 includes a plurality of embeddedmagnet segments 34 for interacting with a levitating magnetic fieldcreated by a levitation magnet structure (not shown) that would bedisposed against or incorporated in housing 22 a. First disc 25 alsocontains embedded magnet segments 35 for magnetically coupling with amagnetic field from a motor (not shown) disposed against or incorporatedin housing 22 b. Housing 22 a includes an inlet 28 for receiving bloodfrom a patient's ventricle and distributing it to vanes 26. Impeller 21is preferably circular and has an outer circumferential edge 30. Byrotatably driving impeller 21 in a pumping direction 31, the bloodreceived at an inner edge of impeller 21 is carried to outercircumferential 30 and enters a volute region 32 within pumping chamber23 at an increased pressure. The pressurized blood flows out from anoutlet 33 formed by housing features 33 a and 33 b. A flow-dividingguide wall 36 may be provided within volute region 32 to help stabilizethe overall flow and the forces acting on impeller 21.

FIG. 3 shows a ventricular pressure curve 40 and a ventricular volumecurve 41 according to a typical cardiac cycle. A pressure pulse 42 and avolume ejection 43 correspond with a systolic phase 44. A lowventricular pressure 45 and increasing ventricular volume 46 correspondwith a diastolic phase 47. The start of systole corresponds with thetime that the mitral or tricuspid valve closes, and the start ofdiastole corresponds with the time that the aortic valve or pulmonaryvalve closes. Curve 48 shows a pulsatile pump flow in which the pumpspeed is synchronously varied in order to provide an increased speedduring systolic phase 44 and a decreased speed during diastolic phase47. As explained above, this conventional pulsatile flow does notsignificantly unload the ventricle at the beginning of systole.

FIGS. 4 and 5 show a first embodiment of the invention for providing avariable speed mode with certain speed changes preceding the beginningof systole 50 and the beginning of diastole 51. An impeller speed curve52 provides a pulsatile flow between an elevated speed 53 and a reducedspeed 54. Curve 49 shows a representative current vector applied to themotor by the motor controller in order to generate a correspondingtarget speed. Thus, an increase in current generates a speed increaseand a decrease in current generates a speed decrease. The target speedincludes a ramping up at segment 55 from reduced speed 54 to elevatedspeed 53, wherein the ramping up begins during diastole. The time ofincrease and the slope of the increase are configured to provide anincreasing flow before systole begins to provide a reduced ventricularpressure that allows a weak ventricle to contract with reducedresistance. The controlled motor current 49 begins to increase at a time56 which precedes systole by a time period t₁. The beginning of rampingup segment 55 of the speed may lag the current increase but still occursprior to the beginning of systole 50. Preferably, time 56 is scheduledby the motor controller at time t₁ before the next expected occurrenceof systole 50 such that the ramping up begins at a moment between about50% to about 90% into the diastolic phase. Thus, denoting the length ofthe diastolic phase as t_(D), the ratio t₁/t_(D) is preferably between0.1 and 0.5.

To help avoid collapse of the ventricle toward the end of systole orduring diastole, impeller speed 52 preferably ramps down at segment 57from elevated speed 53 to reduced speed 54. Segment 57 begins during thesystolic phase of the cardiac cycle (i.e., before the beginning ofdiastole 51). For example, current curve 49 starts to ramp down at atime 58 which precedes start of diastole 51 by a time t₂. Preferably,time 58 may be at a moment between about 50% to about 90% into thesystolic phase. Thus, denoting the length of the systolic phase ast_(s), the ratio t₁/t_(s) is preferably between 0.1 and 0.5.

As shown in FIG. 5, an average speed 59 is maintained as theinstantaneous speed varies between elevated speed 53 and reduced speed54. Average speed 59 may be determined in a conventional manneraccording to the physiological state of the patient. Offsets fromaverage speed 59 for elevated speed 53 and reduced speed 54 may besubstantially constant or may also be determined according to thephysiological state of the patient.

A pump system of the present invention is shown in greater detail inFIG. 6 wherein a controller 60 uses field oriented control (FOC) tosupply a multiphase voltage signal to the pump motor which comprises astator assembly 61 shown as a three-phase stator. Individual phases A,B, and C are driven by an H-bridge inverter 62 functioning as acommutation circuit driven by a pulse width modulator (PWM) circuit 63in controller 60. A current sensing circuit 64 associated with inverter62 measures instantaneous phase current in at least two phases providingcurrent signals designated i_(a) and i_(b). A current calculating block65 receives the two measured currents and calculates a current i_(c)corresponding to the third phase as known in the art. The measuredcurrents are input to an FOC block 66 and to a current observer block 67which estimates the position and speed of the impeller as known in theart. The impeller position and speed are input to FOC block 66.

An average target speed or rpm for operating the pump is provided by aphysiological monitor 68 to FOC block 66. The average rpm may be set bya medical caregiver or may be determined according to an algorithm basedon various patient parameters such heart beat. Monitor 68 may alsogenerate a status signal for identifying whether the ventricle is in theinitial, highly weakened state or whether a predetermined recovery hasbeen obtained in the strength of the ventricle. The average rpm and thestatus signal are provided to a speed command calculator 70. The statussignal can be used to determine whether or not the variable speedcontrol of the invention should be used to unload the ventricle. Thestatus signal can alternatively be externally provided to calculator 70(e.g., by a physician via an HMI).

Command calculator 70 is coupled to a cycle tracking block 71 whichmaintains timing for a cardiac cycle reference. A current signal (e.g.,currents i_(a), i_(b), and i_(c)) can be used in order to detect thecardiac cycle from the instantaneous blood flow, for example. Morespecifically, the controller may identify the heart rate by measuringtime between current peaks in the speed control mode. Then the speeddecrease can start at a calculated time after the occurrence of acurrent peak. The speed increase can start at a calculated time afterthe current minimum value is to detected. This calculated time typicallydepends on the heart rate.

Alternatively, cycle tracking block 71 can be coupled to a pacemaker 72in the event that the patient is using such a device. Conventionalpacemakers have been constructed to continuously generate radio signalsthat contain information about pulse timing and other data. Thesesine-wave modulated signals can be received by a special receiver (notshown), where the signals are demodulated, digitized (if necessary), andtransferred to cycle tracking block 71. Besides being located near theimplanted pacemaker and connected by a cable or wirelessly to thecontroller (e.g., via BlueTooth), a receiver could be integrated withthe controller or the pumping unit.

Based on the reference cycle timing from block 71, command calculator 70determines an instantaneous speed (or magnitude of the current vector)to be used by FOC block 66. FOC block 66 generates commanded voltageoutput values v_(a), v_(b), and v_(c) which are input to PWM block 63.The v_(a), v_(b), and v_(c) commands may also be coupled to observer 67for use in detecting speed and position (not shown). Thus, the speed iscontrolled to follow the curves shown in FIGS. 4 and 5.

In one embodiment, the timing of the speed increases and decreases aredetermined as follows. At a constant pacing rate (i.e., constant beatrate), the time for starting the speed acceleration (e.g., at time 56 inFIG. 4) is:t _(acc)(n+1)=t _(p)(n)₊60/N−t ₁.where t_(p)(n) is the time of occurrence of a pacemaker pulse timesignaling the start of the current cardiac cycle; N is the heart (pulse)rate in beat/min set by a pacemaker; and t_(acc)(n+1) is the time toincrease the pump speed for the next cardiac cycle. Similarly, the timeto start deceleration (e.g., at a time 58 in FIG. 4) is:t _(decel)(n+1)=t _(a)(n+1)+t _(s)where t_(s) is the duration of systole. Systole typically lasts 30% to50% of the cardiac cycle 60/N, and within a certain heart rate range itis fairly independent of the heart rate N. For example, for a heart rateN between 60-120 beats/min, t_(s) is between 0.30 seconds and 0.25seconds.

In an alternative embodiment, command calculator 70 and FOC block 66 areconfigured to operate the motor in a constant current mode (i.e., aconstant torque mode). In this mode, the speed changes inversely withthe pump load (i.e., the flow rate). Thus, an average speed isdetermined by the physiological monitor. The motor controller adjuststhe current to obtain the desired average speed and to keep the currentsubstantially constant. By keeping a constant current in the face of aload which varies within the cardiac cycle, the impeller speedautomatically changes. FIG. 7 shows a load curve 75, wherein the load(i.e., flow rate) is high at 76 during systole and low at 77 duringdiastole. The load ramps up at 78 before the beginning of systole due toan increase of pressure within the ventricle and a decrease of pressureat the pump outlet (e.g., at the aorta). The load ramps down at 80during the beginning of diastole.

In the current control mode, the pump flow increases (load increases) inthe beginning of systole (at 78) and the speed curve 81 drops to areduced speed 83. At the end of systole, the flow drops (at 80) andspeed increases to an elevated speed 82. Thus, the speed increases andstays relatively high during diastole to help unload the ventricle bypumping out blood at the time it fills the ventricle. This is a naturalbehavior of the pump in the current control mode.

Either the variable speed control mode using a variable target speed orusing the constant current approach of the invention can be combinedwith the conventional constant speed mode in order to adapt pumpperformance to the strength level of the patient's ventricle. Inparticular, the selection between the variable speed mode and theconstant speed mode can be determined according to a physiologiccapability of the patient. For example, the pump is set to operate inthe constant current mode immediately following the implantation whenthe left ventricle is weak, thereby providing a greater level ofventricle unloading. With the patient's recovery, the pump may be set tooperate in the constant speed mode, promoting higher flow pulsatilityand a more natural physiologic response to the patient's activities.

What is claimed is:
 1. A blood pump, comprising: a chamber; an impellerdisposed within the chamber; a motor that is configured to drive theimpeller; and at least one processor that: determines a timing of acardio cycle of a user; and varies a speed of the motor based on thetiming of the cardio cycle of the user such that a speed of the motorbegins ramping down during a systolic phase of the cardio cycle.
 2. Theblood pump of claim 1, wherein: varying the speed of the motor comprisesreducing a current level applied to the motor.
 3. The blood pump ofclaim 1, wherein: an average speed of the motor is determined based on aphysiological state of the user.
 4. The blood pump of claim 1, wherein:determining the timing of the cardio cycle comprises measuring a timebetween current peaks.
 5. The blood pump of claim 4, wherein: the speedof the motor begins ramping down starts at a time after one of thecurrent peaks and before a current minimum value is detected.
 6. Theblood pump of claim 1, wherein: determining the timing of the cardiocycle is based on a signal from a pacemaker.
 7. The blood pump of claim1, wherein: the processor varies the speed of the motor based on thetiming of the cardio cycle of the user such that a speed of the motorbegins ramping up during a diastolic phase of the cardio cycle.
 8. Ablood pump, comprising: a chamber; an impeller disposed within thechamber; a motor that is configured to drive the impeller; and at leastone processor that: determines a time period of one or both of adiastolic phase or a time period of a systolic phase of a cardio cycleof a user; and ramps down a speed of the motor during the systolicphase.
 9. The blood pump of claim 8, wherein: the speed of the motor isramped down beginning at a time between about 50% and 90% into thesystolic phase.
 10. The blood pump of claim 8, wherein: the speed of themotor is ramped down to a speed that is based on a physiological stateof the user.
 11. The blood pump of claim 8, wherein: the speed of themotor is ramped down to a substantially constant reduced speed.
 12. Theblood pump of claim 8, wherein: the speed of the motor is ramped down byreducing a current level applied to the motor.
 13. The blood pump ofclaim 8, wherein: the time period of one or both of the diastolic phaseor the time period of the systolic phase is determined by measuring atime between current peaks.
 14. A method for controlling a motor of ablood pump, comprising: determining a timing of a cardio cycle of auser; and varying a speed of the motor of the blood pump based on thetiming of the cardio cycle of the user such that a speed of the motorbegins ramping down during a systolic phase of the cardio cycle.
 15. Themethod for controlling a motor of a blood pump of claim 14, furthercomprising: ramping up a speed of the motor during a diastolic phase ofthe cardio cycle.
 16. The method for controlling a motor of a blood pumpof claim 15, further comprising: maintaining the speed of the motor fora period of time during the diastolic phase prior to ramping up thespeed of the motor.
 17. The method for controlling a motor of a bloodpump of claim 15, wherein: ramping up the speed of the motor begins at amoment between about 50% and 90% into the diastolic phase.
 18. Themethod for controlling a motor of a blood pump of claim 15, wherein: thespeed of the motor completes ramping up during the systolic phase of thecardio cycle.
 19. The method for controlling a motor of a blood pump ofclaim 14, wherein: the speed of the motor completes ramping down duringa diastolic phase of the cardio cycle.
 20. The method for controlling amotor of a blood pump of claim 14, wherein: determining the timing ofthe cardio cycle is based on a signal from a current sensor of themotor.